Signal processor for a hearing device and method for operating a hearing device

ABSTRACT

A signal processor for a hearing device with an implantable stimulator having two or more electrodes for emitting electric charge pulses to neural-fibres of an individual. The processor has a signal path comprising an input circuit adapted to receive an acoustic-signal from the surroundings and provide at least one corresponding input audio-signal; a filter bank adapted to provide atleast one band-limited audio-signal in dependence on the at least one input audio-signal; and a noise filter adapted to attenuate undesired signal components in the at least one band-limited audio-signal and to provide at least one corresponding noise-filtered signal. The processor is characterised in that the portion of the signal path preceding the noise filter neither causes an effective level compression nor an effective level expansion of the atleast one noise-filtered signal when the atleast one noise-filtered signal is derived from an acoustic signal having a level within the comfortable acoustic range.

This application is a Divisional of copending application Ser. No.14/328,228, filed on Jul. 10, 2014, which claims priority under 35U.S.C. §119(a) to Application No. EP13176154.6, filed in the EuropeanPatent Office on Jul. 11, 2013, all of which are hereby expresslyincorporated by reference into the present application.

TECHNICAL FIELD

The present invention relates to a signal processor for a hearing deviceand to a method for operating a hearing device. More specifically, thepresent invention relates to a signal processor for a hearing device forelectric stimulation of nerve cells and to a method for operating such ahearing device.

The invention may e.g. be useful in applications such as a hearing aidfor compensating a hearing-impaired person's loss of hearing capabilityor a listening device for augmenting a normal-hearing person's hearingcapability.

BACKGROUND ART

U.S. Pat. No. 4,207,441 discloses an auditory prosthesis equipmentcomprising n electrode sets implanted in the cochlea at n discretelocations chosen to allow the brain to identify n discrete frequenciesin the sound range. Signals collected by a microphone are passed througha compressor to adapt the dynamic range of the sound information (e.g.60 dB) to the dynamic characteristics of the ear (e.g. 4 dB). Ananalysing network transforms the compressed signal into n analysissignals using filters with frequencies corresponding to the frequenciesidentifiable by the brain. An amplitude calculating network calculatesthe mean value of the amplitudes of each analysis signal, and a signalforming network generates n pulse signals in dependence on the analysissignals. A logic circuit provides a raster signal with width-modulatedpulses in dependence on the n mean values, the n pulse signals and amatching network for preadjustment to individual users. The rastersignal is transmitted inductively to an implanted receiver, whichsequentially transmits electric pulses to respective electrode sets inthe n electrode sets such that the energy transmitted in each electricpulse corresponds to the energy indicated in the respective pulse in theraster signal.

The compression of the microphone signals in the disclosed auditoryprosthesis equipment inherently changes the frequency composition in thesignals, and thus creates artefacts in the n analysis signals providedby the subsequent frequency filtering. Furthermore, the signalprocessing in the disclosed auditory prosthesis equipment does notprovide any specific enhancement of speech signals compared to otheracoustic signals.

SUMMARY OF THE INVENTION

The present invention provides a signal processor that allowsimplementing a hearing device that does not suffer from the aboveproblems. The invention also provides a method for operating a hearingdevice, which overcomes the above problems.

The invention is achieved by the invention defined in the accompanyingindependent claims and as explained in the following description. Theinvention is further achieved by the embodiments defined in thedependent claims and in the detailed description of the invention.

In the present context, a “hearing device” refers to a device, such ase.g. a hearing aid or a listening device, which is adapted to improve,augment and/or protect the hearing capability of a user by receivingacoustic signals from the user's surroundings, generating correspondingaudio signals, possibly modifying the audio signals and providing thepossibly modified audio signals as audible signals to at least one ofthe user's ears in the form of electric signals transferred directly orindirectly to the cochlear nerve, to other sensory nerves and/or to theauditory cortex of the user.

A hearing device may comprise a single unit or several unitscommunicating electronically with each other. Each of the one or moreunits of a hearing device may be configured to be worn in any known way,e.g. behind the ear, entirely or partly arranged in the pinna and/or inthe ear canal, as an entirely or partly implanted unit, etc.

More generally, a hearing device comprises an input transducer forreceiving an acoustic signal from a user's surroundings and providing acorresponding input audio signal, a signal processing circuit forprocessing the input audio signal and an output means for providing anaudible signal to the user in dependence on the processed audio signal.Some hearing devices may comprise multiple input transducers, e.g. forproviding direction-dependent audio signal processing. In some hearingdevices, the output means may comprise one or more output electrodes forproviding electric signals. In some hearing devices, the outputelectrodes may be implanted in the cochlea or on the inside of the skullbone and may be adapted to provide the electric signals to the haircells of the cochlea, to one or more hearing nerves and/or to theauditory cortex.

A “hearing system” refers to a system comprising one or two hearingdevices, and a “binaural hearing system” refers to a system comprisingone or two hearing devices and being adapted to cooperatively provideaudible signals to both of the user's ears. In a hearing system or abinaural hearing system, one or both of the hearing devices may compriseother output means in addition to output electrodes in order to provideaudible signals e.g. in the form of acoustic signals radiated into theuser's outer ears or acoustic signals transferred as mechanicalvibrations to the user's inner ears through the bone structure of theuser's head and/or through parts of the middle ear. In such hearingdevices, the output means may comprise an output transducer, such ase.g. a loudspeaker for providing an air-borne acoustic signal or avibrator for providing a structure-borne or liquid-borne acousticsignal. In a binaural hearing system, the output electrodes may beomitted in one hearing device comprising such other output means.

Hearing systems or binaural hearing systems may further comprise“auxiliary devices”, which communicate with the hearing devices andaffect and/or benefit from the function of the hearing devices.Auxiliary devices may be e.g. remote controls, remote microphones, audiogateway devices, mobile phones, public-address systems, car audiosystems or music players. Hearing devices, hearing systems or binauralhearing systems may e.g. be used for compensating for a hearing-impairedperson's loss of hearing capability, augmenting or protecting anormal-hearing person's hearing capability and/or conveying electronicaudio signals to a person.

As used herein, the singular forms “a”, “an”, and “the” are intended toinclude the plural forms as well (i.e. to have the meaning “at leastone”), unless expressly stated otherwise. It will be further understoodthat the terms “has”, “includes”, “comprises”, “having”, “including”and/or “comprising”, when used in this specification, specify thepresence of stated features, integers, steps, operations, elementsand/or components, but do not preclude the presence or addition of oneor more other features, integers, steps, operations, elements,components and/or groups thereof. It will be understood that when anelement is referred to as being “connected” or “coupled” to anotherelement, it can be directly connected or coupled to the other element,or intervening elements may be present, unless expressly statedotherwise. As used herein, the term “and/or” includes any and allcombinations of one or more of the associated listed items. The steps ofany method disclosed herein do not have to be performed in the exactorder disclosed, unless expressly stated otherwise.

BRIEF DESCRIPTION OF THE DRAWINGS

The invention will be explained in more detail below in connection withpreferred embodiments and with reference to the drawings in which:

FIG. 1 shows an a hearing device according to an embodiment of theinvention,

FIG. 2 shows a functional block diagram of the hearing device shown inFIG. 1 according to an embodiment of the invention,

FIG. 3 shows a diagram illustrating the function of the hearing deviceshown in FIG. 1 according to an embodiment of the invention, and

FIG. 4 shows a diagram illustrating the function of the hearing deviceshown in FIG. 1 according to another embodiment of the invention.

The figures are schematic and simplified for clarity, and they just showdetails, which are essential to the understanding of the invention,while other details are left out. Throughout, like reference numeralsand/or names are used for identical or corresponding parts.

Further scope of applicability of the present invention will becomeapparent from the detailed description given hereinafter. However, itshould be understood that the detailed description and specificexamples, while indicating preferred embodiments of the invention, aregiven by way of illustration only, since various changes andmodifications within the scope of the invention will become apparent tothose skilled in the art from this detailed description.

DETAILED DESCRIPTION OF THE INVENTION

The embodiment of a hearing device 1 shown in FIG. 1 comprises awearable device 2 and an implantable stimulator 3. The wearable device 2comprises a pre-processor 4, a transmission coil 5 and a battery (notshown) for powering the electronic circuits of the wearable device 2and/or the electronic circuits of the implantable stimulator 3. Theimplantable stimulator 3 comprises a reception coil 6, a post-processor7 and a flexible electrode carrier 8. The flexible electrode carrier 8comprises an electrode array 9 with twenty electrodes 10. In otherembodiments, the number of electrodes 10 may be different. Thepre-processor 4 comprises a transmitter 11, and the post-processor 7comprises a corresponding receiver 12.

The wearable device 2 is adapted to be worn on the body of the user ofthe hearing device 1, such that the pre-processor 4 may receive anacoustic signal from the user's surroundings and pre-process theacoustic signal. The transmitter 11 encodes the pre-processed signal andtransmits the encoded signal to the implantable stimulator 3 by means ofthe transmission coil 5. The implantable stimulator 3 is adapted to beimplanted in the body of the user, e.g. on the inside of the skull or inthe cochlea, with the electrodes 10 adjacent to neural fibres such thatelectric charge pulses emitted by the electrodes 10 may stimulate theseneural fibres and thus create a sensation in the user, preferably in theform of a perceived sound. The reception coil 6 is adapted to bearranged such that the post-processor 7 may receive the encoded signalfrom the transmitter 11 by means of the reception coil 6 and thereceiver 12, decode the encoded signal by means of the receiver 12 andprovide electric charge pulses to the neural fibres through theelectrodes 10 of the flexible electrode carrier 8 in dependence on thedecoded signal. The hearing device 1 may thus create sensations in theuser in dependence on the acoustic signal.

The hearing device 1 controls the stimulation, and thereby the sensationin the user, by varying the pulse emission times and/or the amount ofelectric charge emitted in the electric charge pulses. The latter maypreferably be done by varying the duration and/or the amplitude of theelectric charge pulses. Such variations of the charge amount istypically perceived as variations in the strength of the signal, e.g. asloudness in the case that the stimulated neural fibres belong to theauditory nerve. The hearing device 1 is preferably further adapted tostimulate at least two different sets of neural fibres in order tocreate a further dimension in the perception of the sensation. Forexample, stimulating different sets of neural fibres in the auditorynerve may create perception of different sound frequencies in the user.In the following, the term “auditory channel” denotes any distinct setof neural fibres of a particular user that may be stimulated by thehearing device 1.

As shown in the functional block diagram in FIG. 2, the hearing device 1comprises an input circuit 20, a filter bank 21, a noise filter 22, apulse controller 23, a pulse generator 24 and the electrode array 9. Theinput circuit 20 comprises an input transducer 201, a preamplifier 202,a digitiser 203 and a pre-emphasis filter 204. The filter bank 21comprises an FFT converter 211, an energy estimator 212 and a channelcombiner 213. The input circuit 20, the filter bank 21, the noise filter22 and the pulse controller 23 together constitute a signal processor25. In the shown embodiment, the signal processor 25 is comprised by thepre-processor 4. In other embodiments, the signal processor 25 may becomprised partly by the pre-processor 4 and partly by the post-processor7, and in such embodiments the signal processor 25 may further comprisethe pulse generator 24.

The input transducer 201 is arranged such that it may receive anacoustic signal from the user's surroundings and provide a correspondingelectric input signal. The preamplifier 202 amplifies the electric inputsignal, and the digitiser 203 digitises the amplified input signal. Thepre-emphasis filter 204 pre-filters the digitised input signal, suchthat low and high audio frequencies are emphasised in order to achieve afrequency characteristic more like the human ear's natural frequencycharacteristic. The pre-emphasis filter 204 is preferably programmablein order to allow e.g. an audiologist to adapt the frequencycharacteristic to the preferences and the hearing capability of theuser. The input circuit 20 provides the pre-filtered input signal as aninput audio signal to the filter bank 21. In other embodiments, theinput circuit 20 may provide the electric input signal, the amplifiedinput signal or the digitised input signal as the input audio signal. Inother embodiments, the input circuit 20 may comprise further inputtransducers 201, preamplifiers 202, digitisers 203, pre-emphasis filters204, wireless receivers, etc., that allow the hearing device 1 toreceive and process acoustic signals from further sources and/or toachieve directional preference for specific spatial directions.

The FFT converter 211 transforms the input audio signal from atime-domain representation to a frequency-domain representation by meansof a Fast Fourier Transformation (FFT), thus providing a number, e.g.32, 64 or 128, of frequency-bin signals. The energy estimator 212squares the respective frequency-bin signals and provides acorresponding number of frequency-energy signals indicating theinstantaneous energy in the respective frequency-bin signals. Thechannel combiner 213 combines the frequency-energy signals, or anysubset hereof, into a number of frequency channels, each comprising oneband-limited audio signal. The number of frequency channels preferablyequals the maximum number of auditory channels that may be stimulated bythe electrodes 10 in the electrode array 9 and is thus typically lowerthan the number of frequency-bin signals and frequency-energy signals.In other embodiments, the number of frequency channels may be larger, oralternatively, lower than the maximum number of stimulatable auditorychannels. The number of frequency channels may e.g. equal 1, 2, 4, 8,16, 20, 24, 28 or 32. In other embodiments, other well knowntime-to-frequency domain converters may replace the FFT converter 211,such as e.g. a bank of narrow-band filters.

The channel combiner 213 preferably performs the combining for eachfrequency channel by adding one or more of the frequency-energy signalsinto the respective band-limited audio signal. Within any frequencychannel, the channel combiner 213 may weigh the respectivefrequency-energy signals differently, e.g. to accommodate for userpreferences. The channel combiner 213 preferably combines thefrequency-energy signals such that the band-limited audio signalscomprise different subsets of frequency-energy signals. In otherembodiments, one or more subsets of the total set of band-limited audiosignals may comprise identical subsets of frequency-energy signals. Thechannel combiner 213 preferably combines the frequency-energy signalssuch that any frequency-energy signal is at most combined into one ofthe band-limited audio signals. In other embodiments, the samefrequency-energy signal may be combined into two or more of theband-limited audio signals.

The filter bank 21 thus provides at least one band-limited audio signalin dependence on the input audio signal. The filter bank 21 preferablyprovides two or more band-limited audio signals such that the user maydistinguish between acoustic signals having different frequency content.The filter bank 21 is preferably programmable in order to allow e.g. anaudiologist to allocate specific frequency-bin signals to specificband-limited audio signals and thus to specific auditory channelsdepending on the preferences and the hearing capability of the user.

The noise filter 22 attenuates undesired signal components in theband-limited audio signals and provides corresponding noise-filteredsignals. Preferably, one noise-filtered signal is provided for eachfrequency channel and thus for each band-limited audio signal. The FFTtransformation of the input audio signal in the filter bank 21 isperformed using a limited time window in order to keep the time delaythrough the hearing device at a reasonably low level. This inherentlycauses pure tones to smear out on adjacent frequency bands, and sincethe frequency-domain signals are not converted back into time-domainsignals, which would otherwise remove or reduce the smearing, thesmearing is preferably removed or at least reduced by the noise filter22. Several methods for such “side-lobe cleaning” are already known inthe art, and any of these may be implemented in the noise filter 22.

Furthermore, the noise filter 22 preferably removes 50 Hz or 60 Hz noisesignals that may be induced in the transmission coil 5 and/or thereception coil 6 when the coils 5, 6 are close to mains power wiring andthus, depending on the actual configuration of the hearing device 1 (seefurther below), may also appear in the band-limited audio signals.Furthermore, the noise filter 22 preferably removes noise signals with alevel below a noise-floor threshold in order to prevent that the userperceives the noise floor of the surroundings when no sounds of interestare present. The noise filter 22 is preferably programmable in order toallow e.g. an audiologist to adapt the noise attenuation to thepreferences and the hearing capability of the user.

Preferably, the pulse controller 23 causes the pulse generator 24 toprovide one stream of electric charge pulses for each noise-filteredsignal. The pulse controller 23 computes emission times for the electriccharge pulses, preferably such that the emission times coincide with aregular time interval of e.g. 1 ms or 2 ms. In other embodiments,alternative time schemes may be used as is well known in the art. Thepulse controller 23 further computes target charge amounts E (see FIG.3) for the individual electric charge pulses in dependence on thenoise-filtered signals. The computation of the target charge amounts Eis explained in detail in the description of FIG. 3 further below.

The pulse generator 24 generates electric charge pulses and provides theelectric charge pulses to the electrodes 10, such that the emittedelectric charge in each pulse corresponds to the respective targetcharge amount E. The pulse generator 24 preferably provides eachelectric charge pulse as a current flowing out through one or moreelectrodes 10, those electrodes 10 thus having positive polarity, andback through one or more other electrodes 10, those electrodes 10 thushaving negative polarity, thereby causing the current to flow from thepositive electrodes 10 to the negative electrodes 10 through the tissueand thereby stimulating neural fibres in or adjacent to the tissue.

The pulse generator 24 preferably comprises a pulse generator circuit241 and a switching circuit 242 operable to electrically connect thepulse generator circuit 241 to the electrodes 10 of the electrode array9 in different stimulation configurations in order to cause the pulsecurrent to flow through different tissue portions and thereby stimulatedifferent sets of neural fibres. The pulse generator 24 may thusstimulate different auditory channels and create different sensations inthe user by changing the stimulation configuration of the pulsegenerator 24 and/or the electrode array 9. Instead of, or in additionto, switching the electrodes 10, multiple, individually controllablepulse generator circuits 241 may be used to achieve the same or furtherconfigurability.

In one or more stimulation configurations, a first electrode 10 may havepositive polarity and a second, adjacent electrode 10 may have negativepolarity while the remaining electrodes 10 are disconnected from thepulse generator circuit 241 and thus have neutral polarity, such thatthe pulse current flows from the first to the second electrode 10. Inone or more other stimulation configurations, a first electrode 10 mayhave positive or negative polarity while all of the remaining electrodes10 have respectively negative or positive polarity, such that the pulsecurrent flows respectively from the first electrode 10 to all of theremaining electrodes 10 or in the opposite direction. In one or morefurther stimulation configurations, a first electrode 10 and allelectrodes 10 between the first electrode 10 and the respective end ofthe electrode array 9 may have positive polarity while a second,adjacent electrode 10 and all electrodes 10 between the second electrode10 and the respective other end of the electrode array 9 may havenegative polarity, such that the pulse current flows substantiallybetween two portions of the electrode array 9, which together compriseall of the electrodes 10. Further stimulation configurations, such asmixtures of the above stimulation configurations, may be used ifdesired.

Preferably, each frequency channel or pulse stream is allocated to apredefined auditory channel, and the predefined auditory channels andtheir allocation are preferably chosen such that sensations perceived bythe user are similar to the sensations a normal-hearing person wouldperceive when subjected to acoustic signals within the frequency rangesof the corresponding band-limited audio signals.

The pulse generator 24 preferably provides the electric charge pulsessuch that they each start at a respective computed emission time. Thecharge amount provided by each pulse may be controlled by varying theduration and/or the amplitude of the electric charge pulses. The pulsegenerator 24 preferably provides the electric charge pulses as bi-phasepulses, i.e. as pairs of pulses with the two pulses of each pairfollowing immediately after each other and having opposite polarity andthus opposite current direction, in order to prevent a build-up overtime of electric potential in the tissue. In this case, the two pulsesof each bi-phase pulse preferably each provides one half of therespective target charge amount E.

The pre-processor 4 preferably comprises the signal processor 25, andthus the input circuit 20, the filter bank 21, the noise filter 22 andthe pulse controller 23, as well as a transmitter 11. The pulsecontroller 23 provides a pulse signal indicating the computed emissiontimes and the computed target charge amounts E for the individualelectric charge pulses. The transmitter 11 receives the pulse signal asthe pre-processed signal, encodes the pulse signal and transmits theencoded pulse signal to the implantable stimulator 3 through thetransmission coil 5. The transmission is preferably performed usingnear-field magnetic induction (NFMI) signals that are known to penetrateskin and tissue without significant attenuation. The encoded pulsesignal is preferably transmitted by means of amplitude modulation, e.g.on/off keying, of a carrier signal with a frequency in the low MHzrange, e.g. about 1-10 MHz or about 2-5 MHz. Accordingly, thepost-processor 7 preferably comprises a receiver 12 and the pulsegenerator 24. The receiver 12 receives the encoded pulse signal throughthe reception coil 6, decodes the encoded pulse signal into the pulsesignal and provides the decoded pulse signal to the pulse generator 24.The pulse generator 24 generates the electric charge pulses inaccordance with the emission times and target charge amounts E indicatedin the decoded pulse signal as described further above.

The hearing-device configuration described above allows for performing aminimum of power-consuming computations and/or signal processing withinthe post-processor 7 and thus allows for supplying power to thepost-processor 7 using the NFMI signals transmitted by the transmitter11, e.g. by using these signals to charge a capacitor (not shown)comprised by the implantable stimulator 3. The transmission is performedusing a transmission protocol that preferably allows for transmittingdummy signals in order to provide further power to the post-processor 7in cases and/or during time periods wherein the power transmitted in theencoded pulse signal does not suffice. Other hearing-deviceconfigurations may be readily contemplated. However, the benefits ofsuch other hearing-device configurations should be balanced against thepossibly increased complexity of the circuits to be implanted, which maymake the implantable stimulator 3 less robust and further increase itspower consumption. Also, the amount of data to be transmitted to theimplantable stimulator 3 may increase, which may require an increase intransmission bandwidth and further increase the complexity of theimplanted receiver 12. In such other configurations, the transmitter 11and the receiver 12 should obviously appear in the appropriate otherfunctional blocks of the functional diagram.

The diagram shown in FIG. 3 illustrates how target charge amounts E forthe individual electric charge pulses in each pulse stream or auditorychannel may be computed from the levels L of the respectivenoise-filtered signal using a piecewise linear mapping function, e.g. asimple mapping function 30 or an enhanced mapping function 31. The Laxis and the E axis are both linear, and each mapping function 30, 31maps the logarithm L of the sound pressure or the energy in therespective noise-filtered signal into a target charge amount E for therespective pulse stream or auditory channel. The mapping functions 30,31 may differ between the auditory channels due to different couplingbetween the electrodes 10 and the neural fibres and/or differentsensitivity in the different neural fibres.

A normal-hearing person has a frequency-dependent hearing threshold thatdefines the weakest sounds the person can hear and a frequency-dependentuncomfortable level (UCL) that defines the weakest sounds that causediscomfort in the person. In the following text, the term “comfortableacoustic range” refers to the frequency-dependent level range betweenthe typical hearing threshold and the typical UCL for normal-hearingpersons. Statistically obtained values for these levels are well knownin the art. It is also well known that normal-hearing persons typicallyperceive loudness approximately logarithmic within the comfortableacoustic range. A specific relative increase, such as e.g. a doubling,of the sound pressure or the energy in an acoustic signal thus typicallycreates a perception of a specific absolute increase of loudness. Thisis reflected in the common use of the logarithmic dB scale for soundlevels.

In contrast hereto, the perception of loudness when the auditory neuralfibres are directly electrically stimulated is approximately linearwithin the dynamic range perceivable by the user. A specific absoluteincrease in the emitted electric charge thus typically creates aperception of a specific absolute increase of loudness, regardless ofthe starting loudness. The perception is, however, quite varying fromperson to person, and the hearing device 1 therefore needs to becalibrated to the individual user. After the implantation of theimplantable stimulator 3, an audiologist thus normally performs one ormore tests in order to determine which electrodes 10 are functional,which auditory channels may be stimulated and which stimulation levelsmay be perceived by the user in the individual auditory channels. Foreach auditory channel, the audiologist preferably determines a thresholdcharge T equal to the smallest charge that the user can perceive incharge pulses and a maximum comfortable charge C equal to the largestcharge that the user can perceive in charge pulses emitted repeatedlyfor a longer time without feeling discomfort. In the following text, theterm “comfortable stimulation range” refers to the level range betweenthe threshold charge T and the maximum comfortable charge C for aparticular auditory channel.

The input circuit 20, the filter bank 21 and the noise filter 22together constitute a signal path taking the acoustic signal as inputand providing the noise-filtered signal(s) as output. The portion of thesignal path 20, 21, 22 preceding the noise filter 22 preferably appliesa linear gain, i.e. a gain that is not level-dependent, to any acousticsignal within the comfortable acoustic range. Thus, the signal path 20,21, 22 preceding the noise filter 22 neither causes an effective levelcompression nor an effective level expansion of noise-filtered signalsderived from an acoustic signal having a level within the comfortableacoustic range. This preferably also applies to the noise filter 22.Electronic components used for implementing the signal path 20, 21, 22are thus preferably dimensioned such that their inherent nonlinearitiesneither affect the band-limited audio signal nor the band-limited audiosignal noise-filtered signals when these signals are derived from anacoustic signal having a level within the comfortable acoustic range.Alternatively, or additionally, the signal path 20, 21, 22 may comprisemeans for compensating for such nonlinearities in order to avoid botheffective level compression and effective level expansion.

Also, the input gain, i.e. the total gain in the signal path 20, 21, 22,is preferably fixed or calibrated at known values. This allows thehearing device 1 to let the user perceive the loudness of an acousticsignal similarly strong as a normal-hearing person by mapping thecomfortable acoustic range into the comfortable stimulation range suchthat an acoustic signal at the hearing threshold causes stimulation atthe threshold charge T and an acoustic signal at the UCL causesstimulation at the maximum comfortable charge C in the respectiveauditory channel with a linear dependency between these end points. Whenthe input gain is known, the noise signal level L may be computed as thesum of the level of the acoustic signal and the input gain, at least fora tonal acoustic signal that is not attenuated by the noise-filter 22,and thus, the linear mapping from the comfortable acoustic range intothe comfortable stimulation range may be achieved by mapping acorresponding level range of the noise signals into target charge amountE, e.g. using the simple mapping function 30.

In each point (L, E), the mapping functions 30, 31 have a slope dE/dL,which for instance in the example point (L₁, E₁) determines how much thetarget charge amount E shall increase from E₁ when the noise-filteredsignal level L increases from L₁. In the following, the slope dE/dL isreferred to as the incremental gain G_(i). The sound levels L providedto the mapping functions 30, 31 may attain negative, zero and/orpositive values depending on the reference value for the signal levelsL. On the other hand, the electric charge amounts E provided by themapping functions 30, 31 are inherently non-negative. Note therefore,that an absolute gain is not well defined and that the unit of theincremental gain G_(i) obviously depends on the respective units of thesound levels L and of the target charge amounts E.

The simple mapping function 30 preferably extends linearly between alower knee point 32 and an upper knee point 33. In the lower knee point32, the simple mapping function 30 maps a threshold level L_(T)corresponding to the user's hearing threshold into the threshold chargeT. In the upper knee point 33, the simple mapping function 30 maps amaximum comfortable level L_(C) corresponding to the user's UCL into themaximum comfortable charge C. Between the lower and upper knee points32, 33, the simple mapping function 30 has a constant and positiveincremental gain G_(i) of (C-T)/(L_(C)-L_(T)). Between the knee points32, 33, the electric charge amounts emitted by the electrodes 10 thusincrease with increasing noise-filtered signal levels L and thus withincreasing levels of the acoustic signal. Signal levels L above themaximum comfortable level L_(C) are preferably mapped into the maximumcomfortable charge C and thus with zero incremental gain G_(i) in orderto avoid creating uncomfortable sensations in the user. Signal levels Lbelow the threshold level L_(T) are preferably mapped into zero charge Cand thus with infinite incremental gain G_(i) in order to save energy inthe implantable stimulator 3. The noise-floor threshold of the noisefilter 22 is preferably above and not below the threshold level L_(T).The noise-floor threshold of the noise filter 22 may be set equal to thethreshold level L_(T) in order to allow the user to perceive acousticsignal levels within the entire comfortable acoustic range.

The computation of target charge amounts E using a piecewise linearmapping function, such as the simple mapping function 30, effectivelyamounts to applying a level-dependent gain in the hearing device 1. Suchlevel compression and/or level expansion are commonly applied and thuswell known in the art of hearing devices. However, by applying levelcompression and/or level expansion only after deriving the band-limitedaudio signals in the filter bank 21 and after attenuating undesiredsignal components in the noise filter 22, thus allowing the filter bank21 and the noise filter 22 to operate on signals that have neither beensubjected to effective level compression nor to effective levelexpansion, the amount of artefacts produced by the filter bank 21 andthe noise filter 22 is substantially reduced. Furthermore, by applyinglevel compression and/or level expansion only after deriving theband-limited audio signals in the filter bank 21 and after attenuatingundesired signal components in the noise filter 22, parameters for levelcompression and/or level expansion may be set in a more deterministicway than in the prior art, i.e. in dependence on predefined levels ofthe received acoustic signal. Thus, predefined acoustic levels may bemapped to predefined stimulus levels, i.e. pulse charge amounts, whichallows an audiologist to perform fitting of the hearing device 1 in amuch more intuitive way than possible for prior art hearing devices.

The enhanced mapping function 31 deviates from the simple mappingfunction 30 only in that it extends from the lower knee point 32 to theupper knee point 33 via an intermediate knee point 34, such that theincremental gain G_(i) is larger between the lower knee point 32 and theintermediate knee point 34 than between the intermediate knee point 34and the upper knee point 33. In the intermediate knee point 34, theenhanced mapping function 31 maps an intermediate threshold level L_(K)to an intermediate charge K. Compared to the simple mapping function 30,the enhanced mapping function 31 thus applies a level expansion toaudible signal levels L below the intermediate threshold level L_(K) anda level compression to comfortable signal levels L above theintermediate threshold level L_(K).

The purpose of the intermediate knee point 34 is primarily to enhanceinformation conveyed in speech and thus improve the user's ability todecode and understand speech. During typical conversations with moderatevoice levels in quiet surroundings, the level of the speech signalsreceived by the hearing device 1 normally concentrate within the lowerportion of the comfortable acoustic range. The intermediate thresholdlevel L_(K) is therefore preferably determined in dependence on astatistic evaluation of typical speech situations such that the majorityof the levels typically appearing in such speech situations is below theintermediate threshold level L_(K). The influence of the input gain onthe noise-filtered signal levels L should be appropriately considered inthis determination. The intermediate charge K is preferably closer tothe maximum comfortable charge C than to the threshold charge T. Thesetwo conditions on the intermediate knee point 34 are normally sufficientto place it above the curve of the simple mapping function 30. However,for some auditory channels and/or for some speech languages, they maynot suffice. In this case, the simple mapping function 30 should be usedinstead of the enhanced mapping function 31 for the particular auditorychannel(s). The intermediate knee point 34 ensures that the levelstypically appearing in speech situations are mapped to a relativelylarger range of target charge amounts E, which typically will improvethe user's ability to understand the speech.

The enhanced mapping function 31 may be used also in such hearingdevices 1 and/or signal processors 25 wherein the absence of effectivelevel compression and effective level expansion applies only to asmaller portion than stated further above of the level range of theacoustic signal—or does not apply at all, even when this may increasethe amount of artefacts. In such hearing devices 1 and/or signalprocessors 25, the location of the knee points 32, 33, 34 in the diagramof FIG. 3 may have to be adjusted to best fit the thus possiblylevel-dependent gain in the signal path 20, 21, 22.

In other embodiments, the intermediate threshold level L_(K) may bedetermined such that about 60%, about 70%, about 80% or about 90% of thelevels typically appearing in the speech situations are below theintermediate threshold level L_(K). The determination of theintermediate threshold level L_(K) may further be made based onpreferences of the user and/or on knowledge about the user's hearingability in general. Similarly, the intermediate charge K may bedetermined such that it is located at a distance from the thresholdcharge T equalling about 60%, about 70%, about 80% or about 90% of thedistance between the threshold charge T and the maximum comfortablecharge C. Also the determination of the intermediate charge K mayfurther be based on preferences of the user and/or on knowledge aboutthe user's hearing ability in general. For speech at a medium level inthe French language, and for the average user, the intermediate kneepoint 34 may preferably be determined according to Table 1 below. Thefrequency ranges and the intermediate knee points 34 may vary, e.g. forother languages and for different users.

TABLE 1 Frequency range L_(K)-(input gain) (K-T)/(C-T) 200 Hz-850 Hz 61dB_(SPL) 90%  850 Hz-1500 Hz 61 dB_(SPL) 90% 1500 Hz-3450 Hz 57 dB_(SPL)90% 3450 Hz-8000 Hz 50 dB_(SPL) 90%

Table 1 should be interpreted such that for noise-filtered signalsmainly comprising frequencies within the frequency range 200 Hz-850 Hz,the intermediate threshold level L_(K) should thus be set to 61 dBSPLplus the input gain for the respective frequency range, and theintermediate charge K should be set at 90% of the distance from thethreshold charge T to the maximum comfortable charge C, etc. The kneepoints 34 defined by Table 1 are set such that about 90% of theinformation comprised in the listed speech frequency bands are below therespective intermediate threshold levels L_(K).

As described earlier; for each auditory channel, the threshold charge Tequal to the smallest charge that the user can perceive in charge pulsesis determined. The target charge amounts E for individual electriccharge pulses in each pulse stream or auditory channel is mapped to thelevels L of respective noise-filtered signal. For example, the thresholdcharge T is mapped to the threshold level L_(T) of the comfortableacoustic range. Conventionally, signals lower than this minimum level donot stimulate the patient while the signal above stimulate the patient.In an embodiment (as represented in FIG. 4), for overcomingnon-linearity on the behavior of the simulation that may be considerablefor signals that are in average around the threshold level L_(T),simulation below the threshold level is provided. This would bebeneficial to the patient to the patient as the nerves would be on nearsimulation threshold when a signal level L is close to the thresholdlevel L_(T). As illustrated, the mapping functions 30, 31 is extended ina region where the noise-filtered signal level L is below the thresholdlevel L_(T). For the enhanced mapping function 31, such extension ismade utilizing the slope dE/dL obtained between the lower knee point 32and the intermediate knee point 34. Although, the extended simplemapping function 30 may be utilized for detemining the stimuation butthe extended enhanced mapping function is preferred. In such a scenario,the stimulation may be limited by either the minimum system sensitivity(MSS) or the absolute minimum capability (MC) of the implant such as 5μs.

The hearing device 1 may preferably comprise a user-operable control(not shown), such as a control element on the wearable device 2 and/oron a wired or wireless remote control, that allows the user to adjustthe intermediate threshold level L_(K). The user-operable controlpreferably allows the user to adjust the intermediate threshold levelL_(K) in predefined level steps, such as e.g. 10 dB steps or 6 dB steps.This allows the user to adapt the signal processing in the signalprocessor 25 to weaker and/or louder speech. The signal processor 25 maypreferably comprise an environment analyser (not shown) that analysesthe acoustic signal and adjusts the intermediate threshold level L_(K)in dependence on the analysis, such that the intermediate thresholdlevel L_(K) decreases when the signal processor 25 receives weakerspeech signals and increases when the signal processor 25 receiveslouder speech signals. The signal processor 25 may thus automaticallyperform adjustments otherwise made by the user.

The hearing device 1 may preferably comprise a user-operable control(not shown) that allows the user to switch the hearing device betweenusing the simple mapping function 30 and the enhanced mapping function31 for relevant auditory channels, i.e. those auditory channels havingan intermediate knee point 34 above the simple mapping function 30. Thisallows the user to adapt the signal processing in the hearing device 1to situations with and without speech. Thus, when speech is present, theuser may choose signal processing than enhances speech, and when speechis absent, the user may choose signal processing that provides a morenatural loudness curve, e.g. for environmental sounds. The signalprocessor 25 may preferably comprise a speech detector that detectsspeech in the acoustic signal and switches the signal processor 25 tousing the enhanced mapping function 31 for the relevant auditorychannels when detecting speech and switches the signal processor 25 tousing the simple mapping function 30 when speech is absent, thusautomatically performing switching otherwise made by the user.

In some embodiments, the used mapping functions 30, 31 may have furtherknee points. In some embodiments, the mapping functions 30, 31 may abovethe upper knee point 33 have a positive incremental gain G_(i) that issmaller than the incremental gain G_(i) between the lower and the upperknee points 32, 33. In some embodiments, the mapping functions 30, 31may have an incremental gain G_(i) that is zero or positive and finitebelow the lower knee point 32. The knee points 32, 33, 34 are preferablyimplemented as soft knee points, i.e. such that the incremental gainG_(i) transitions smoothly in the immediate vicinity of the knee points32, 33, 34, in order to avoid abrupt changes of the incremental gainG_(i) in the knee points 32, 33, 34. In some embodiments, theincremental gain G_(i) may be only approximately constant, and themapping functions 30, 31 may be only approximately piecewise linear.

The signal processor 25 is preferably implemented mainly as digitalcircuits operating in the discrete time domain, but any or all partshereof may alternatively be implemented as analog circuits operating inthe continuous time domain. Digital functional blocks of the hearingdevice 1, such as e.g. the pre-emphasis filter 204, the filter bank 21,the noise filter 22, the pulse controller 23, the transmitter 11, thereceiver 12 and the pulse generator 24 may be implemented in anysuitable combination of hardware, firmware and software and/or in anysuitable combination of hardware units. Furthermore, any single hardwareunit may execute the operations of several functional blocks in parallelor in interleaved sequence and/or in any suitable combination thereof.

The hearing device 1 may be part of a binaural hearing system.

Further modifications obvious to the skilled person may be made to thedisclosed apparatus and/or method without deviating from the scope ofthe invention. Within this description, any such modifications arementioned in a non-limiting way.

Some preferred embodiments have been described in the foregoing, but itshould be stressed that the invention is not limited to these, but maybe embodied in other ways within the subject-matter defined in thefollowing claims. For example, the features of the described embodimentsmay be combined arbitrarily, e.g. in order to adapt the system, theapparatus and/or the method according to the invention to specificrequirements.

It is further intended that the structural features of the apparatusdescribed above, in the detailed description of ‘mode(s) for carryingout the invention’ and in the claims can be combined with the methods,when appropriately substituted by a corresponding process. Embodimentsof the methods have the same advantages as the corresponding apparatus.

Any reference numerals and names in the claims are intended to benon-limiting for their scope.

1. A hearing aid system comprising an implantable stimulator configuredto emit electric charge pulses to neural fibres of an individual; andthe signal processor circuit including an input circuit configured toreceive a signal from individual's surroundings and provide at least onecorresponding input audio signal, a filter bank configured to receivethe at least one input audio signal and to output at least oneband-limited audio signal in dependence on the at least one input audiosignal, and a noise filter configured to attenuate undesired signalcomponents in the at least one band-limited audio signal and to provideat least one corresponding noise-filtered signal; and a pulse controllerconfigured to apply a level dependent compression and/ or expansion onthe at least one noise filtered signal by computation of target chargeamounts in accordance with a mapping function.
 2. The hearing aid systemaccording to claim 1, wherein the input circuit and the filter bank areconfigured to prevent an effective level compression and an effectivelevel expansion of the at least one noise filtered signal when the inputcircuit and the filter bank process the signal having a level within theindividual's predefined comfortable acoustic range.
 3. The hearing aidsystem according to claim 1, wherein the pulse controller is configuredto control the emission of the electric charge pulses for at least oneauditory channel in dependence on the at least one noise-filteredsignal, such that the amount of electric charge provided in the electriccharge pulses increases with increasing level (L) of the at least onenoise-filtered signal, thereby defining at least one level-dependentincremental gain.
 4. The hearing aid system according to claim 1,wherein the pulse controller is further configured to control theincremental gain such that it is larger for any level of the at leastone noise-filtered signal between a lower threshold level and an upperthreshold level than for any level above the upper threshold level. 5.The hearing aid system according to claim 1, wherein the pulsecontroller is further configured to control the incremental gain suchthat it is larger for any level of the at least one noise-filteredsignal between the lower threshold level and an intermediate thresholdlevel than for any level between the intermediate threshold level andthe upper threshold level.
 6. The hearing aid system according to claim1, further comprising an environment analyser configured to analyse thesignal in the individual's environment and adjust the intermediatethreshold level in dependence on the analysis, such that theintermediate threshold level decreases when the signal processorreceives weaker speech signals and increases when the signal processorreceives louder speech signals.
 7. The hearing aid system according toclaim 5, wherein the signal processor is configured to control theintermediate threshold level in dependence on input provided by theindividual.
 8. The hearing aid system according to claim 4, wherein thelevel thresholds are determined to correspond with respectivepredetermined threshold levels of the signal.
 9. The hearing aid systemaccording to claim 1, wherein the pulse controller is configured tocontrol the incremental gain such that it increases linearly orpiecewise linearly for any level of the at least one noise-filteredsignal between the lower level threshold and the upper level threshold,except for smooth transitions in the immediate vicinity of the levelthresholds.
 10. The hearing aid system according to claim 1, wherein theinput circuit comprises: an input transducer configured to receive thesignal and provide a corresponding electric input signal; and adigitiser configured to digitise the electric input signal and toprovide the digitised signal as the input audiosignal.
 11. The hearingaid system according to claim 10, wherein the input circuit furthercomprises: a pre-amplifier configured to apply a linear gain to theelectric input signal, the gain being independent of the level of theelectric input signal; and a pre-emphasis filter that is configuredbased on a hearing capability of the individual to achieve a human ear'snatural frequency characteristic.
 12. The hearing aid system accordingto claim 1, wherein the filter bank comprises: a time-to-frequencydomain converter configured to provide a set of narrow-band frequencysignals in dependence on the input audio signal; and a channel combinerconfigured to provide the at least one band-limited signal as a sum oran energy sum of a subset of the set of narrow-band frequency signals.13. The hearing aid system according to claim 1, wherein the implantablestimulator comprises two or more electrodes for emitting electric chargepulses to neural fibres of an individual.
 14. The hearing aid systemaccording to claim 1, wherein the input circuit is configured to receivethe signal comprising an acoustic signal from individual's surroundingsand provide the at least one corresponding input audio signal.
 15. Awearable device comprising a hearing aid system according to claim 1.16. A method for operating a hearing device comprising an implantablestimulator having two or more electrodes for emitting electrical chargepulses to neural fibres of an individual, the method comprising:receiving asignal and providing a corresponding input audio signal;inputting the input audio signal to a filter bank and outputting atleast one band-limited audio signal in dependence on the input audiosignal; inputting the at least one band-limited audio signal to a noisefilter and attenuating undesired signal components in the at least oneband-limited audio signal and providing at least one correspondingnoise-filtered signal; and computing a target charge amounts inaccordance with a mapping function and applying a level dependentcompression and/ or expansion on the at least one correspondingnoise-filtered signal in accordance with the computed target chargeamounts.
 17. The method according to claim 14, further comprisescontrolling the emission of the electric charge pulses for at least oneauditory channel in dependence on the at least one noise-filteredsignal, such that the amount of electric charge provided in the electriccharge pulses increases with increasing level of the at least onenoise-filtered signal, thereby defining at least one level-dependentincremental gain.
 18. The method according to claim 14, furthercomprising preventing an effective level compression and an effectivelevel expansion of at least one band-limited audio signal prior to theattenuating of the undesired signal components by the noise filter. 19.The method according to claim 14, further comprising controlling theincremental gain such that it is larger for any level of the at leastone noise-filtered signal between a lower threshold level and an upperthreshold level than for any level above the upper threshold level. 20.The method according to claim 14, further comprising controlling theincremental gain such that it is larger for any level of the at leastone noise-filtered signal between the lower threshold level and anintermediate threshold level than for any level between the intermediatethreshold level and the upper threshold level.